The invention is directed to biodegradable, thermoplastic, multi-block copolymers. The copolymers of the present invention find use in various applications, particularly in the field of pharmaceutical drug delivery systems, drug-eluting coatings and biomedical implants.
The invention relates to biodegradable multi-block copolymers, the hydrolysable sequences being amorphous and the segments being linked by a multifunctional chain-extender, and the segments having different physical and degradation characteristics. For example, a multi-block co-polyester consisting of a glycolide-ε-caprolactone segment and a lactide-glycolide segment is composed of two different polyester pre-polymers. By controlling the segment monomer composition, segment ratio and length, a variety of polymers with properties that can easily be tuned can be obtained. These materials are particularly interesting for constructing drug delivery matrices, which contain and release a therapeutic agent, such as injectable drug-loaded biodegradable microspheres for controlled drug delivery, or drug-eluting coatings for medical devices.
Considerable research has been undertaken in the field of drug delivery matrices that contain and deliver various biologically active agents.
One reason for these research efforts is to develop pharmaceutical delivery systems, which prolong the release time of existing drugs. Many new drugs have short half-lives, which necessitates frequent injection schedules. Another reason is that many new drugs that may have been developed have poor pharmacokinetic profiles. In particular, peptides and proteins cause pharmacokinetic difficulties. Such substances must be administered parenterally if systemic action is required. Patient compliance and the high costs associated with frequent dosing protocols for parenterally administered drugs provide strong stimuli for the development of alternative dosage forms and dosing regimens.
Delivery matrices provided in the form of a coating, e.g. on a medical device, are often referred to as drug-eluting coatings. The major driver for the use of drug-eluting coatings is to improve the performance of the medical device, i.e. more successfully treating the disease and/or preventing or reducing undesired side reactions, such as inflammation or infection. Drug-eluting coatings allow the controlled release of biologically or pharmacologically active compounds due to which a therapeutically effective drug concentration can be achieved over a certain period of time. Drug-eluting coatings further allow local site-specific drug delivery. The drug can be delivered locally thereby allowing the achievement of high concentrations of the active compound at the site where it is most needed. Total drug doses may be significantly lowered thereby preventing the high systemic concentrations associated with oral administration of the frequently highly toxic drugs.
Polymeric systems which are presently under investigation as biodegradable drug delivery matrices for injectable or implantable pharmaceutical formulations and as drug-eluting coatings to be applied on medical devices include poly-D,L-Lactide (PDLLA), copolymers of lactide and glycolide (PLGA) (Brannon-Peppas, Int. J. Pharmaceutics, 116 (1995) p1-9; Couvreur, et al., Advanced Drug Delivery Reviews, 10 (1993) p141-162; Conti, et al., J. Microencapsulation, 9 (1992) p153-166 and copolymers of lactide and ε-caprolactone (Buntner et al, J. Control. Rel. 56 (1998) 159). PGLA is by far the most widely applied matrix system for injectable drug delivery systems. In the field of biodegradable materials for drug-eluting coatings, Drachman et al. (J. of American College of Cardiology, vol. 36, no. 7, 2000) reported on the use of poly(lactide-ε-caprolactone) (PLA-ε-CL) copolyesters as biodegradable coating material for the controlled release of paclitaxel from vascular stents. A fully degradable heparin-eluting (PLA-ε-CL) stent has been reported by Gao R. et al. (J. of American College of Cardiology, vol. 27, no. 85A, 1996 (abstract)). Polylactide (DL-PLA) and a polylactide-trimethylenecarbonate copolymer (PLA-co-TMC) have been used for the controlled release of dexamethasone from Strecker stents (Strecker E. P., et al., Effect on intimal hyperplasia of dexamethasone released from coated metal stents compared with non-coated stents in canine femoral arteries, Cardiovasc. Intervent. Radiol., 21 1998 p. 487. EP1254674 describes a polylactide acid (Mw=30 kDa) based stent coating for the controlled local delivery of tacrolimos. Bertrand O. F. et al., (Biocompatibility aspects of new stent technology, J. of American College of Cardiology, vol. 32, no. 3, 1998) reviewed several materials for use as matrix material for a drug-eluting coating. Van der Giessen et al. (Marked inflammatory sequalae to implantation of biodegradable and non-biodegradable polymers in porcine coronary arteries. Circulation 94 (1996) 1690) evaluated several materials, including PGLA and PCL for application as a drug-eluting coating on stents. Prietzel et al (Inhibition of neointimal proliferation with a novel hirudin/prostacyclin analog eluting stent coating in an animal overstretch model. Circulation 94 (1996) I-260) and Lincoff et al. (sustained local delivery of dexamethasone by a novel intravascular eluting stent to prevent restinosis in the porcine coronary injury model. J. Am. Coll. Cardiol. 29 (1997) 808-816) tested PLLA as a matrix for controlled delivery of their active compounds.
Amorphous PLGA copolymers and PDLLA homopolymers have a number of disadvantages when used in controlled drug-release applications. Due to their high sub-body temperature Tg's, both PLGA and PDLLA are rigid matrices. The ability to manipulate the release of an encapsulated drug, especially if the drug has a high molecular weight such as proteins, is therefore limited because of a limited diffusion of these molecules within PLGA and PDLLA matrices. The release of drugs from PLGA and PDLLA matrices, therefore, is initially solely governed by diffusion of dissolved drug molecules through pores. Only in a later stage, when hydrolytic degradation has lowered the molecular weight sufficiently or when (parts of) the polymer matrix start to dissolve, diffusion of drug molecules through the polymer matrix becomes possible, generally leading to dose dumping of the encapsulated drug. Furthermore, during degradation of PLGA and PDLLA, acidic degradation products (lactic and glycolic acid) are accumulating in the polymeric matrix due to its glassy character (Tg >37° C.), which may have a negative effect on sensitive actives such as proteins and peptides, but may be harmless to other drugs. Random copolymers of lactide and caprolactone (PLA-ε-CL) yield less acidic degradation products. Moreover, these copolymers are not associated with significant pH reductions in the polymer matrix if the polymer matrix is rubbery under body conditions, i.e. the Tg (glass transition temperature) of the copolymer is below appr. 37° C. Under these conditions, the polymer matrix is also permeable to high molecular weight drugs and to the degradation products that are released, thereby preventing accumulation and as a result preventing the generation of an acidic environment. However, these materials are very sticky due to which processing into free flowing microspheres, which is a typical prerequisite for the formulation of injectable particulate drug delivery systems, is rather challenging. For the same reason, they are also difficult to handle when used as drug-eluting coatings on medical devices. Sticking can be greatly reduced by increasing the lactide content, but then the polymer will become too rigid. Increasing the caprolactone content can also reduce sticking, but then the overall degradation rate of the polymers becomes so low that accumulation of the polymer material at the site of the injection might occur upon repeated injections.
Furthermore, the physicochemical properties of the above mentioned (co)polymers can only be affected by three parameters: molecular weight, monomer ratio and monomer distribution, which is an important drawback if optimization of characteristics of pharmaceutical or medical formulations is required. Because the reactivity of glycolide, lactide and caprolactone towards ring-opening is very different and the high temperatures that are usually required for complete monomer conversion, it is difficult to obtain a controlled monomer distribution in this type of copolymers. Therefore, also randomly polymerized terpolymers, which are built of these monomers are not suitable enough to create a matrix of polymers with a wide range of polymer properties. Thus, there is an obvious need for the provision of new materials for drug delivery applications that overcome the above mentioned disadvantages of the currently used materials and provide better tools to control and optimize the characteristics of pharmaceutical or medical formulations, especially with respect to the release characteristics of encapsulated drugs.
This can be achieved by the use of biodegradable multi-block copolymers of the present invention, comprising segments of pre-polymers of different chemical composition and physico-chemical characteristics. By combining different segments with different physico-chemical properties, different functionalities can be built into the material, e.g. high swelling degree, increased permeability, or slow degradation rate. Moreover, weak and disadvantageous properties of one of the segments may be masked, whereas advantageous properties of the individual segments may be combined. Moreover, a functionality may be introduced without directly affecting other functionalities of the polymer. Moreover, by combining two segments of different composition, leading to a certain extent of phase separation, biphasic release patterns can be achieved. For example, if one of the segments has low permeability and/or is slowly degrading and the other segment has high permeability and/or degrades rapidly, encapsulated drug molecules will initially be released predominantly from the phase which has a high permeability and/or degrades rapidly, before release of drug molecules encapsulated in the phase with lower permeability/degradation rate will start to contribute significantly. By modifying the permeability and/or degradation rates of the two phases, the time at which release from a specific phase starts can be controlled.
Copolymers used as pharmaceutical drug delivery matrices, such as injectable microspheres or drug-eluting coatings, do not necessarily need to be rigid under body conditions. It is even considered an advantage that a drug-eluting implant, coating or microsphere is soft as this prevents tissue irritation due to mechanical friction. It is however beneficial that the materials are rigid under processing conditions as to prevent sticking. The latter is especially relevant when a therapeutical agent is encapsulated in microspheres if one wants to collect them as individual particles, and re-suspension of the (freeze-)dried formulation prior to injection may be problematic due to agglomerate formation. More important, injection into the body will be problematic, as the needle may be blocked easily due to agglomeration of the microspheres.
Besides the previously mentioned amorphous homo- and copolymers, also many block co-polyesters (AB, ABA and multi-block) have been studied in the past and are still under investigation for their drug loading and release properties. ABA type block copolymers that comprise a hydrophilic, non-biodegradable block such as polyethylene glycol (PEG) and a hydrolysable polyester block intended for drug release purposes are described in patent application U.S. Pat. No. 5,548,035. These copolymers are built of a polyethylene glycol central block and hydrophobic hydrolysable non-swellable outer hard block consisting of PLA, PGA, PLGA or PCL. Also block copolymers based on amorphous ester blocks (A) and hydrophilic ether groups B have been prepared. The amorphous character of the group A improves the solubility in organic solvents compared to ABA blocks based on crystallisable PGA or PLA sequences. ABA block (PELA) copolymers comprising poly(D,L-Lactide) (A) and PEG (B) blocks have also been studied for drug and protein loading efficiency: (Deng X M, Li X H, Yuan M L et al., J. Control. Release 1999, 58 123-31: Optimization of preparative conditions for poly DL-Lactide-polyethylene glycol microspheres with entrapped Vibrio cholera antigens” Deng X M, Zhou S B, Li X H, Zhao J, Yuan M L. In vitro degradation and release profiles for poly DL-Lactide-polyethylene glycol microspheres containing human serum albumin. J. Control. Release 2001, 71, 165-73. Other examples are disclosed in patent number U.S. Pat. No. 6,258,121, which describes the use of a blend of a hydrophilic lactide-polyethylene oxide copolymer (PLA-PEO) and a hydrophobic PLA-ε-CL copolymer as a stent coating for the controlled local delivery of paclitaxel into the blood vessel wall. Although polymer properties can be greatly improved by using block copolymers with blocks of different copolymers instead of homo- or random copolymers, they still have some disadvantages.
To obtain a minimum molecular weight of a block copolymer with e.g. an ABA structure, the sequences A and B must have a certain length. The blocks may independently behave as the individual homopolymers with similar composition. Properties of the ABA type block co-polymers can mainly be tuned by varying the composition of A and B blocks or, alternatively, by varying the A/B ratio or the length of blocks A and B. The properties of multi-block co-polymers wherein the blocks or segments are much shorter and linked together by a chemical reaction can also be affected by varying segment length and ratio. Properties such as stiffness, permeability, swelling characteristics and degradation behaviour and also drug release characteristics can be tuned in a much better way. Examples of known multi-block co-polymers are those described by Penco et al. (M. Penco, S. Marcioni, P. Ferruti, S. D'antone and R. Deghenghi, Biomaterials 17 (1996) 1583-1590) concerning multi-block co-polymers containing poly(lactide/glycolide) and polyethylene glycol segments and by Li et al. (S. Li, H. Gareau. M. Vert, T. Petrova, N. Manolova, I. Rashkov, J. of Appl. Pol. Science 68, (1998) 989-998 who describes multi-block co-polymers of poly(ε-caprolactone) and polyethylene glycol segments. However, both type of multi-block co-polymers contain only one hydrolysable polyester segment, the other segment being non-hydrolysable (PEG). The freedom to vary with degradation and physical properties is therefore mainly restricted to the composition of the hydrolysable segment.
Penco et al. (European Polymer Journal, 36 (5), 2000, 901-908) also studied and described the preparation of amorphous multi-block co-polymers with structure (PGLA50/50-PCL530)n comprising two different hydrolysable amorphous segments. However, by using their preparation method, comprising phosgene, only multi-block co-polymers with alternating PCL530 en PGLA50/50 segments can be obtained. This method is therefore restricted to the preparation of multi-block co-polyesters with equimolar amounts of the two individual segments, thus limiting the possibilities to vary the composition and monomer distribution of multi-block co-polymers.